It thus constitutes a peristaltic type pump which is designed so that from the physiological point of view the behavior of the present pump is close to that of nature so as to avoid any hemolysis phenomenon (i.e. destruction of red corpuscles), and to reduce as far as possible shock phenomena due to sudden changes of pressure in the blood fluid. Insofar as possible, the design of the pump must also ensure that its operation is essentially non-displacement in order to avoid the risk of venous collapse.
Several systems have already been proposed for making a heart ventricular prosthesis.
Of the most recent implantable systems, mention may be made initially of the system including a closed flat bag having each of its two walls in contact with a respective presser plate for the purpose of compressing the bag to expel blood fluid therefrom. These two plates are hinged to the free ends of two levers that move apart from or towards each other like scissors blades, the opposite ends of the two levers being hinged on a common support and each lever supporting a respective solenoid in the vicinity of said common support. That assembly is described in particular in an article entitled "Implantable LVDA" by P. M. Portuer et. col. at pp. 115-141 of "Assisted Circulation--2" edited by Felix Unger, and published in 1984 by Springer-Verlag--Berlin, Heidelberg, New York, Tokyo.
The assembly constituted by the closed bag and the drive means for compressing the bag is disposed inside an essentially rigid outer shell serving both to make the implant comfortable and to protect the components of the prosthesis.
However, the use of a rigid and sealed outer shell gives rise to disturbances in the peristaltic operation of the pump, which disturbances can have consequences that are very severe for the patient since they run the risk of giving rise to venous collapse.
This results from the fact that the displacements of the volumes of blood in the closed bag give rise to pressure variations inside the rigid and sealed shell, such that the underpressure at the end of the ejection stage (delivering blood fluid) may reach a very high value.
For example, with a pump having a shell of 300 cm.sup.3 and a closed bag capable of ejecting 40 cm.sup.3 of blood fluid per systole, the bag is subjected at the end of the ejection stage to underpressure exceeding 70 mm of mercury (i.e. 9,100 Pa which is comparable to the ejection pressure), and which can even reach 100 mm of mercury (i.e. 13,000 Pa), which is a very high value if it is compared with the mean ejection pressure from a natural heart which is about 150 mm of mercury (i.e. 19,500 Pa).
In addition, such a system, whose operation is, substantially of the displacement type, does not make it possible to avoid this physical phenomenon of underpressure, unless, of course, a shell is provided that is of very large volume, but under such circumstances the pump would take up too much room.
U.S. Pat. No. 4,976,729 (Holfert et al.) and U.S. Pat. No. 4,750,903 (Cheng) describe pneumatic systems whose operation is disturbed little or not at all by underpressure in the shell. The underpressure is completely masked by the compression gas or the suction vacuum driving the membrane or the bag. In contrast, the driving energy requirements are increased to compensate the removal of blood. Such an increase is no problem in systems driven from outside the patient, which systems in any case present very low energy efficiency.
The Applicant has proposed an implantable system with non-displacement operation as described in detail in European patent No. 0 148 661.
That system comprises an implantable blood pump of the type comprising an outer shell, at least one essentially flat and deformable closed bag disposed inside the shell, having one of its walls in contact with the inside surface of said shell, and being connected to an inlet valve to enable it to be filled with blood fluid and to an outlet valve to enable it to deliver blood fluid, a cylindrical drum having an end edge of the bag fixed to a generator line thereof and organized to roll over the other wall of said bag, a C-shaped bracket mounted to rotate about the axis of said drum and having its central portion connected by means of a flexible inextensible component to the other end edge of the bag, and an electric motor integrated within the drum having its stator secured to said drum and having its rotor driving said bracket, said motor enabling the bracket and the drum to be driven in mutual and antagonistic rotation through a fraction of a turn to compress the bag which winds onto said drum in order to expel the blood fluid.
The outer shell of the implantable pump described in the above-specified European patent is made of biocompatible material (e.g. silicone rubber or polyurethane) making it possible to associate the geometrically deformable shape of the shell with a material that has its own memory (using a material having its own geometrical memory improves energy restitution since the memory of the material is added to the deformation that results from the relative displacement of the drum and the C-shaped bracket).
Although the use of a flexible outer shell does indeed avoid the above-described phenomenon of underpressure, it suffers in practice from certain drawbacks: the movements of the motorized components are not always tolerated well by the adjacent organs, which constitutes a source of discomfort for the patient; in addition, even with perfect biocompatibility, in the long term rigid new tissue is observed to form on the outside of the shell, giving rise to increasing sensitivity to underpressure during the ejection stage. Furthermore, if the outer shell is too flexible, it may collapse during the movements of the motor-driven components, and in the limit it may jam the bracket.
It would naturally be most advantageous to be able to conserve the structural organization of that pump (in order to retain its non-displacement operation) while using an outer shell that is rigid and hermetically closed (in order to both ensure that the implant is comfortable and to protect the components of the prosthesis).
However, the underpressure physical phenomenon mentioned above for the preceding system is then encountered and it gives rise to a suction force on the closed bag during the ejection stage.
Whatever type of implantable system is used, there are numerous consequences of the underpressure phenomenon:
a considerable reduction in the volume of ejected blood fluid;
a drop in pressure towards the aorta; and
too much energy is consumed by the drive system which must overcome the additional force.
If this physical phenomenon is to be countered, it is necessary to provide the implantable pump with a compensation system.
Several compensation systems may then be considered, and some have indeed been experimented with: the compensation may be rigid and internal, flexible and internal, or it may be external.
However, the following explanations show that such systems do not really give satisfaction.
Known rigid internal compensation systems include rated springs or electromagnets with pistons for moving at the appropriate moment a volume that enables the underpressure in the shell to be compensated.
Systems of this type have been experimented with, but they are not really workable at present. In addition to their large bulk, they suffer from the drawback of requiring mechanical or pneumatic coupling between the compensation volume and means for measuring of the underpressure in the shell, all of which must be placed in the patient's chest. In addition, such systems are not free from danger since the slightest tendency towards positive pressure (overcompensation) prevents the bag filling.
Known flexible internal compensation systems are based on the principle of a special in-body flexible bag whose volume is necessarily large. Such systems are not suitable since the special bag rapidly becomes the seat of rigid new tissue formation: the special compensation bag then progressively looses its ability to deform, and ends up by becoming an additional dead volume like the shell. In any event, the real compensation that could be expected is highly limited from the beginning because of the equilibrium that is established quickly (the system behaves like a closed system).
There then remains the solution of external compensation which remains theoretically possible: if the sealed shell remains at atmospheric pressure via a tube enabling it to "breathe", then system performance returns to that of an open shell. However, this solution requires the use of a transcutaneous catheter which constitutes a path for infection, particularly if the implant is for the long term. In addition, if the implant is completely internal, including its source of energy, such a catheter quickly becomes unacceptable to the patient.
In a variant, attempts have been made to use a catheter with a syringe for creating a small amount of underpressure, but in practice such a solution is hardly any more satisfactory.
Finally, known compensation systems do not make it possible to counter the physical underpressure phenomenon in satisfactory manner when a rigid and sealed shell is used, such that present solutions remain limited to comprises between these systems without real progress.
It is of interest to observe that most specialists, when faced with the need to provide a compensation system, have sought to reduce the dead volume of the shell. However, it appears that such an approach is not suitable since if the volume of the shell is halved, then the underpressure is substantially doubled during ejection, such that the force applied on the motor system becomes very large and all of the unrestrained portions of the flexible bag start to inflate, thereby reducing the ejected volume correspondingly. It is clear that it is impossible to avoid having such unrestrained portions, particularly where the flexible bag is connected to the inlet and outlet valves. It is perhaps possible to reduce the unrestrained surfaces of the flexible bag with a pump having presser plates of the above-described type, but the forces acting on the presser plates rapidly reach several kilograms, thereby preventing the pump from operating unless very large amounts of energy are supplied.
An object of the invention is to provide a higher-performance implantable blood pump whose structure makes it possible to obtain non-displacement pump operation without requiring an additional compensation system.